Radiation detector and X-ray CT apparatus

ABSTRACT

A radiation detector includes a plurality of detector modules detachably mounted on a detector base. Each of the detector modules has a plurality of element blocks permanently mounted on a module base. Each element block has a plurality of radiation detection elements formed on a signal substrate in the form of an m×n matrix. A detector module is made up of a plurality of element blocks. A radiation detector is made up of a plurality of detector modules. This makes it possible to tile many detection elements and manufacture a radiation detector with a wide field of view.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is based upon and claims the benefit of priority fromthe prior Japanese Patent Applications No. 11-366180, filed Dec. 24,1999; and No. 11-368273, filed Dec. 24, 1999, the entire contents ofwhich are incorporated herein by reference.

BACKGROUND OF THE INVENTION

The present invention relates to a 2D array type radiation detectorhaving a plurality of detection elements in the form of a matrix whichdetect radiations such as x-rays as electrical signals, and an x-ray CTapparatus.

A medical x-ray CT apparatus has an x-ray tube and detector. X-raysgenerated by the x-ray tube are transmitted through an object to beexamined and incident on the detector. The detector has a plurality ofdetection elements for detecting radiations such as x-rays as electricalsignals. Detection elements can be classified into indirectionconversion type elements, each designed to convert an x-ray into lightby a phosphor (scintillator) and further convert the light into anelectrical signal by a photoelectric conversion element (photodiode),and direct conversion type apparatuses, each using specificsemiconductor characteristic, i.e., a photoconduction phenomenon inwhich electron-hole pairs are generated in a semiconductor and moved toits electrode by using x-rays. It is expected that direct conversiontype apparatuses, which can achieve reductions in size, weight, andprofile, will become popular.

As detectors for x-ray CT, single-slice type detectors are widely used.A single-slice type detector has a plurality of detection elementsarrayed in a line. A multislice type detector constituted bysingle-slice type detectors arranged in a plurality of lines is alsoknown.

FIG. 1 is a partial sectional view of a conventional multislice typedetector. FIG. 2 is a schematic plan view of the detector. Referring toFIG. 2, an illustration of a scintillator is omitted. A plurality ofphotodiodes 92 are arranged on the rear surface of a scintillator 97.The plurality of photodiodes 92 are respectively connected to aplurality of integrators 95 through a plurality of wires 91. Selectionswitches 96 are provided in units of lines. Outputs from the integrators95 are sequentially read out through the selection switches 96. Theoutputs of the selection switches 96 are electrically connected to asubstrate 94 through bonding wires 93.

The integrators 95 store the signals detected by the photodiodes 92.Integral signals are sequentially output to the substrate 94 by theselection switches 96 through the bonding wires 93. The reason why theintegral signals are sequentially read out by the selection switches 96is that the number of bonding wires that can be formed on the substrate94 is limited.

A great deal of attention has been paid to a 2D array type detector as anext-generation detector, which has more channels than the abovemultislice type detector, with the element pitch in the verticaldirection (slice direction) being equal to the element pitch in thehorizontal direction (channel direction).

To put this 2D array type detector into practice, various problems mustbe solved.

First, as the number of detection elements greatly increases as in the2D array type detector, the precision in tiling the many elements into aspecific shape deteriorates.

Second, as the number of detection elements greatly increases as in the2D array type detector, the probability of the occurrence of faultydetection elements increases, and hence the yield decreases.

Likewise, as the detector is used for a long period of time, it isinevitable that some of many detection elements will fail. In this case,a detection element array or the overall detector must be replaced,resulting in a high cost. This is the third problem.

In addition, signal sampling is performed in CT an enormous number oftimes, e.g., several hundred or thousand times, per rotation. Therefore,the time permitted for 1-period signal read operation is very short. Itis very difficult to complete reads of signals from many channels withinsuch a short period of time. This is the fourth problem.

BRIEF SUMMARY OF THE INVENTION

It is an object of the present invention to provide a radiation detectorwhich implements tiling of many detection elements in the form of amatrix.

A radiation detector includes a plurality of detector modules detachablymounted on a detector base. Each of the detector modules has a pluralityof element blocks permanently mounted on a module base. Each elementblock has a plurality of radiation detection elements formed on a signalsubstrate in the form of an m×n matrix. A detector module is made up ofa plurality of element blocks. A radiation detector is made up of aplurality of detector modules. This makes it possible to tile manydetection elements and manufacture a radiation detector with a widefield of view.

Additional objects and advantages of the invention will be set forth inthe description which follows, and in part will be obvious from thedescription, or may be learned by practice of the invention. The objectsand advantages of the invention may be realized and obtained by means ofthe instrumentalities and combinations particularly pointed outhereinafter.

BRIEF DESCRIPTION OF THE SEVERAL VIEWS OF THE DRAWING

The accompanying drawings, which are incorporated in and constitute apart of the specification, illustrate presently preferred embodiments ofthe invention, and together with the general description given above andthe detailed description of the preferred embodiments given below, serveto explain the principles of the invention.

FIG. 1 is a partial sectional view of a detector in the prior art;

FIG. 2 is a view showing the arrangement of bonding wires connected tothe detector in the prior art;

FIG. 3 is a system diagram of an x-ray CT apparatus according to thefirst embodiment of the present invention;

FIG. 4 is a plan view showing the schematic structure of a radiationdetector 127 in FIG. 3;

FIG. 5 is a view showing the structure of an element block 15 in FIG. 4;

FIG. 6 is a partial sectional view of the element block 15 in FIG. 4;

FIG. 7 is a partial sectional view showing another shape of a notchedportion in FIG. 6;

FIG. 8 is a partial sectional view showing still another shape of thenotched portion in FIG. 6;

FIG. 9A is a side view showing an array of element blocks in the firstembodiment;

FIG. 9B is a side view showing another array of element blocks in thefirst embodiment;

FIG. 9C is a perspective view showing an array of detector modules inthe first embodiment;

FIG. 10A is a view showing the side surface structure of a detectormodule in the first embodiment;

FIG. 10B is a side view showing an array of the detection modules shownin FIG. 10A;

FIG. 11 is a sectional view showing a substrate and its peripheralportion in FIG. 10A;

FIG. 12A is a view showing another side surface structure of thedetector module in the first embodiment;

FIG. 12B is a view showing the side surface structure of a detectormodule paired with the detector module in FIG. 12A;

FIG. 12C is a side view showing an array of the detector modules in FIG.12A and the detector modules in FIG. 12B;

FIG. 13 is a side view showing a grid substituting a collimator;

FIG. 14 is a view showing an example of how a plurality of collimatormodules are mounted in the first embodiment;

FIG. 15 is a cross-sectional view of a detector according to the firstembodiment;

FIG. 16 is a side view showing another scintillator piece in the firstembodiment;

FIG. 17 is a side view showing still another scintillator piece in thefirst embodiment;

FIG. 18 is an equivalent circuit diagram of a detection element in thefirst embodiment;

FIG. 19 is another equivalent circuit diagram of a detection element inthe first embodiment;

FIG. 20 is still another equivalent circuit diagram of a detectionelement in the first embodiment;

FIG. 21 is an equivalent circuit diagram of one detector module in thefirst embodiment;

FIG. 22 is a view showing a signal read sequence for one line ofdetector modules in FIG. 21;

FIG. 23 is another equivalent circuit diagram of one detector module inthe first embodiment;

FIG. 24 is a view showing a signal read sequence for one line ofdetector modules in FIG. 23;

FIG. 25 is still another equivalent circuit diagram of one detectormodule in the first embodiment;

FIG. 26 is a view showing the arrangement of an x-ray CT scanneraccording to the second embodiment of the present invention;

FIG. 27 is a perspective view of a radiation detector in FIG. 26;

FIG. 28 is a view showing processing and the flow of data in the secondembodiment;

FIG. 29 is a view showing an example of image display in the secondembodiment;

FIG. 30 is another example of image display in the second embodiment;and

FIG. 31 is a view showing CT fluoroscopic operation in the secondembodiment.

DETAILED DESCRIPTION OF THE INVENTION

Preferred embodiments of the present invention will be described indetail below with reference to the views of the accompanying drawing.

(First Embodiment)

FIG. 3 is a system diagram of an x-ray CT apparatus according to thefirst embodiment.

An x-ray tube 131 is supported, together with a radiation detector 127,to be rotatable around an object 132 to be examined. The x-ray tube 131generates a so-called x-ray cone beam spreading in two directions,namely a channel direction C and a slice direction A (direction parallelto the rotation axis (direction perpendicular to the drawing surface)).The x-ray beam transmitted through the object 132 is detected by theradiation detector 127. The signal detected by the radiation detector127 is sent to a data processing unit 135 for performing correctionprocessing and the like through a data acquisition circuit 134 toundergo predetermined signal processing. The resultant data istemporarily stored in a storing unit 136. The following components areconnected to a host controller 138: a high-voltage generator 139 forsupplying power to the x-ray tube 131, a gantry driving unit 140 for arotating gantry that rotates the x-ray tube 131 and the like, areconstructing unit 137 for reconstructing data, a display unit 141 fordisplaying the image reconstructed by the reconstructing unit 137, anoperation unit 142 for operating the display unit 141, an input device143 for sending a control signal from the operation unit 142 to the hostcontroller 138, and the like.

FIG. 4 schematically shows the structure of the radiation detector 127.The radiation detector 127 is comprised of a plurality of, e.g., 38,detector modules 34 arrayed along the channel direction C. In x-ray CT,the 38 detector modules 34 are not arrayed flat but are arrayed in theform of an arc centered on the focal point of the x-ray tube 131. Onedetector module 34 is made up of one element module 26 and onecollimator module 33. Each element module 26 is made up of a pluralityof, e.g., four, element blocks 15 arrayed along the slice direction A.One element block 15 has m×n detection elements in the form of a matrixformed on a single substrate, together with peripheral circuits. In thiscase, one detection element is handled as one channel. Obviously,however, a predetermined number of neighboring detection elements may behandled as one channel. The number of channels per block is set inaccordance with, for example, a matrix size of 24×64, which exhibitsrelatively high yield in manufacturing semiconductor devices.

In the block manufacturing stage, the element blocks 15 are inspectedone by one, and defective products are eliminated. A plurality of, e.g.,four, element blocks 15 are arrayed along the slice direction A andfixed on a module base (an element 18 in FIG. 10A). Note that the fourcoupled element blocks 15 will be referred to as the element module 26.The collimator module 33 is mounted on the element module 26, thuscompleting the detector module 34. The element modules 15 can not bedisassembled. The detector is assembled, tested, repaired, and replacedin units of detector modules 34.

The 38 detector modules 34 are arrayed on a curved detector base 28(FIG. 10B), thus completing the radiation detector 127. Each of the 38detector modules 34 is detachably mounted on the detector base. If,therefore, a given detector module 34 fails, the radiation detector 127can be inexpensively and quickly restored by replacing only the faultydetector module with a normal detector module 34.

Note that the radiation detector 127 may be formed by arraying theelement blocks 15 in two orthogonal directions, namely the directions Aand C, without using the element modules 26. However, this detector ispreferably handled in units of modules 34 in consideration of operationefficiency and yield.

FIG. 5 is an exploded perspective view of the element block 15. FIG. 6is a sectional view of this block. Photodiodes 17 are formed of an m×nmatrix, and mounted on the upper surface of a substrate 14. Ascintillator block 16 is mounted on the photodiodes 17. The scintillatorblock 16 is made up of m×n scintillator pieces 11 equal in number to thematrix of photodiodes 17.

A side surface and x-ray incident surface of each scintillator piece 11are coated with a light reflecting material. The light reflectingmaterial blocks external light and prevents leakage of light generatedby each scintillator piece 11. In place of the light reflecting materialcoat, white plastic plates may be bonded to the side surface and x-rayincident surface of each scintillator piece 11.

Most of the scintillator pieces 11 have rectangular parallellepipedshapes, typically cubic shapes. As shown in FIG. 6, however, nscintillator pieces 11 located on the two ends in the slice direction Ahave end faces each notched, obliquely and inwardly, from itssubstantially middle point to the bottom surface so as to have asubstantially pentagonal cross-section. A bonding wire 13 for connectingthe photodiode 17 to the substrate 14 is accommodated in the spacesecured by a notched portion 30. With this structure, as shown in FIGS.9A and 9B, when the four element blocks 15 are joined to each otheralong the slice direction A to form the element module 26, thescintillator pieces 11 of the adjacent element blocks 15 can be broughtinto tight contact with each other, thus eliminating any gaps betweenthe blocks. In addition, since each bonding wire 13 extends from an endportion in the slice direction A, when the detector modules 34 arearrayed in the channel direction C, all gaps between the modules can beeliminated, as shown in FIG. 9C.

Note that the shape of each notched portion 30 is not specificallylimited. For example, as shown in FIG. 7, an end face of thescintillator piece 11 may be obliquely notched from the upper surface tothe bottom surface. In this case, the scintillator piece 11 on thecorresponding end has a trapezoidal cross-section. Alternatively, an endface of the scintillator piece 11 may be notched in a proper curveinstead of being notched straight, as shown in FIG. 8.

The surface area of the element block 15 (substrate 14) is designed tobe almost equal to the x-ray incident surface area of the scintillatorblock 16. The scintillator blocks 16 are designed to have almost thesame size. Note that the size of the photodiodes 17 located on the twoends in the slice direction A may be designed to be slightly smallerthan the size of the remaining photodiodes 17 in consideration of ajoint margin. In this case, the channels at the ends of each elementmodule 26 in the slice direction A tend to greatly differ in x-rayconversion ratio from the remaining channels. However, this problem canbe solved by causing a data processing unit 35 to perform datacorrection such as weighted interpolation for the data detected by thechannels at the ends. Weights are set in consideration of the purpose ofinspection, the precision of data obtained by the elements at the ends,expected resolution, and the like.

The signal detected by each photodiode 17 is sent as an electricalsignal to the substrate 14 through the bonding wire 13. Owing toproblems in boding techniques, the bonding wire 13 protrudes from thesurface of a photodiode 12 to some extent. The protruding bonding wire13 is accommodated in the space defined by the notched portions 30 ofthe two element blocks 15 adjacent to each other in the slice directionA.

FIG. 10A is a side view showing one detector module 34 when viewed fromthe slice direction. As described above, one detector module 34 iscomprised of one element module 26 made up of the four element blocks 15coupled to each other in the slice direction and one collimator module33 mounted on the element module 26. The element module 26 is fixed on aplate-like module base 18 through a fixing stand 31. A data acquisitioncircuit board 25 for reading out signals from the photodiodes 17 andacquiring signals is placed on the element module 26, which is securedby the poles 31. The signal sent from each photodiode 17 to thesubstrate 14 through the bonding wire 13 is sent to the data acquisitioncircuit board 25 of a data acquisition unit 143 placed on the lowersurface of the board through an interconnection in the board, as shownin FIG. 11. Note that this circuit 25 may be formed on the substrate 14of the photodiode 17, together with a photodiode array and itsperipheral circuit.

The collimator module 33 has a plurality of collimator plates 20 eachmade of a heavy metal with high stiffness such as tungsten ormolybdenum. The plurality of collimator plates 20 are supported betweentwo collimator supports 21 to be arrayed parallel at intervals equal tothe pitch of channels. The collimator module 33 is positioned withrespect to the element module 26 such that the plurality of collimatorplates 20 are respectively positioned on the boundaries between aplurality of channels.

The width of the collimator module 33 in the channel direction C isdesigned to be almost equal to that of the element module 26. Thecollimator module 33 is not aligned with the element module 26 but isshifted from the element module 26 in the channel direction C by adistance (Δd/2) ½ a distance (pitch) Ad between the central points ofthe adjacent detection elements (channels). By shifting the collimatormodule 33 from the element module 26 by the distance (Δd/2), thecollimator plate 20 can be positioned immediately above the boundarybetween the channels of the adjacent element modules 26. With thisarrangement, when the 38 detector modules 34 are arrayed in a line on adetector base 28. The collimator plate 20B can be positioned between achannel CA on the right end of a given module 34A and a channel CB onthe left end of an adjacent module 34B on the right. Thus the scatterdradiation removment can be implemented at the boundary.

By sequentially placing the modules 34 having the same structure in thechannel direction C in this manner, the gaps between the modules 34 canbe eliminated.

As described above, each of the 38 detector modules 34 can be easilydetached from the detector base 28 by relatively easy operation, e.g.,unfastening a few screws. With this arrangement, when a given detectormodule 34 fails, the faulty module is detached from the detector base28, and a new normal module 34 is mounted in the empty space, therebyrestoring the normal function of the radiation detector 127.

Note that when the faulty module 34 is to be replaced, since thecollimator module 33 overlaps the adjacent modules 34, the faulty module34 cannot be detached alone, a plurality of normal modules 34 on theright side of the faulty module 34 must also be detached.

FIGS. 12A, 12B, and 12C show a modification configured to improve theefficiency of replacing operation by decreasing the number of detectormodules 34 to be detached when the faulty module 34 is to be replacedwith a normal module 34. For this purpose, two types of detector modules34-1 and 34-2 are prepared. The two types of detector modules 34-1 and34-2 have the same structure except for the widths of collimator modules33-1 and 33-2 and the numbers of collimator plates 20. As shown in FIG.12A, in one collimator module 33-1, the number of collimator plates 20is larger than the number of channels (n) by one. As shown in FIG. 12B,in the other collimator module 33-2, the number of collimator plates 20is smaller than the number of channels (n) by one. One collimator module33-1 is wider than the other collimator module 33-2 by a widthcorresponding to the difference (two) between the numbers of collimatorplates 20.

Two types of detector modules 34 whose collimator modules 33-1 and 33-2differ in this manner are alternately arranged on the detector base 28without any gap along the channel direction C, as shown in FIG. 12C.

This structure requires two types of detector modules 34. However, whena faulty module is to be replaced with a normal module 34, the number ofdetector modules 34 to be detached can be decreased to one or three.When the faulty module 34 in FIG. 12A is to be replaced, only the faultymodule 34 is detached, and a normal module 34 is attached. When thefaulty module 34 in FIG. 12B is to be replaced, the two adjacent modules34 on the two sides of the faulty module 34 are detached, together withthe faulty module 34, and a normal module 34 is attached. Thereafter,the two adjacent modules 34 are placed back into position.

Note that a grid may be used in place of a collimator. FIG. 13 shows anarrangement using a grid. FIG. 13 is an enlarged view of a portion nearthe grid and a scintillator. A grid 24 is formed by alternately stackingand bonding metal foils 22 made of a heavy metal such as lead andintermediate members 23 made of a light metal such as aluminum. Sincethe metal foil 22 is supported by the intermediate member 23, anysupports like the collimator supports 2 are not required.

Note that collimators may be completed by arraying the detector modules34 on which the collimator modules 33 are mounted. As shown in FIG. 14,after the detector modules 34 on which no collimator modules 33 aremounted are arrayed, the collimator modules 33 may be mounted on thedetector modules 34. Alternatively, collimators completed by couplingthe collimator modules 33 may be mounted on the arrayed detector modules34.

As described above, by notching portions of the scintillator pieces 11on the ends and placing extraction means such as the bonding wires 13 inthe notched portions, a large radiation detector without any gap can beformed, which is required to, for example, require temporally continuousvoxel data.

For example, in the prior art, only four channels can be arrayed in theslice direction A. As shown in FIG. 15, according to the presentinvention, 256 channels can be implemented in the slice direction byarraying four element blocks 15 each having m (e.g., 64) photodiodes inthe slice direction A. More channels can be implemented by increasingthe number of element blocks 15 arrayed or arraying a plurality ofelement modules 26 along the slice direction A.

In the overall detector, M×N (256×912) channels can be implemented byarraying four element blocks 15, each having m×n (64×24) photodiodes, inthe slice direction A, and 38 element blocks 15 in the channel directionC. Note that m may be an even number, e.g., m=64, or may be an oddnumber, e.g., m=65. The number represented by m is not limited to aspecific value. In addition, the number of element blocks in the channeldirection C is may be an even number, e.g., n=24, or an odd number,e.g., n=25. Similar to m, the number represented by n is not limited toa specific value. Likewise, the number of detector modules is notlimited to an even or odd number.

If the number of element blocks 15 in the slice direction A is an evennumber, e.g., four as shown in FIG. 9B, the center line of an x-ray beamgenerated by the x-ray tube 131 in the slice direction A passes throughthe joint portion between the element block 15 and another elementblock. If the number of element blocks 15 in the slice direction A is anodd number, e.g., three as shown in FIG. 9C, the center line of an x-raybeam generated by the x-ray tube 131 passes through the center of theelement block 15.

According to the above description, each scintillator piece has arectangular parallelepiped shape. However, as shown in FIG. 16, ascintillator piece 51 having a substantially parallelogrammiccross-section whose upper side on the x-ray incident surface side isslightly longer than the lower side on the light output surface side maybe used, or a scintillator piece having a trapezoidal cross-sectionwhose light output surface is narrower than the x-ray incident surfacemay be used. In addition, a photodiode 52 is positioned and shaped tooppose the light output surface of each scintillator piece 51. Sinceother arrangements are the same as in the first embodiment, adescription thereof will be omitted.

In this case, since the scintillator pieces other than those on the twoends also have shapes other than rectangular parallelepiped shapes, themethod of manufacturing a scintillator block is complicated. However,the light incident surfaces of the photodiodes joined to thescintillator pieces on the two ends of the scintillator block can bemade almost equal in size to those of the photodiodes joined to thescintillator pieces other than those on the two ends, and hence theprecision of data detected at the two ends can be improved.Alternatively, a plurality of scintillator pieces may be selected fromthose on the ends, and each selected scintillator piece may have a shapewhose x-ray incident surface is narrow than the light output surface.

In this case, since the scintillator pieces and photodiodes other thanthose on the two ends of the element block change in shape, the x-rayconversion efficiency may greatly vary. In this case, therefore, dataprecision can be improved by performing data correction such as weightedinterpolation for the data detected by all the scintillator photodiodesas well in the data processing unit 35. Weights should be set inconsideration of the purpose of an inspection, the precision of dataobtained by the elements on the ends, expected resolution, and the like.

Furthermore, in the arrangement shown in FIG. 17, each scintillatorpiece 61 has a substantially rectangular parallelepiped shape whosex-ray incident surface is almost equal in size to the light outputsurface as described above, but a dummy scintillator 62 formed on eachend is thinner than the remaining scintillator pieces. The x-rayincident surface of each dummy scintillator 62 is almost equal in sizeto the x-ray incident surface of the scintillator piece 61, but thelength of a side surface of the dummy scintillator 62 is shorter thanthat of the scintillator piece 61. The dummy scintillator 62 is formedto, for example, shield the bonding wire 13 against x-rays so as toprevent a malfunction. As the dummy scintillator 62, a generalscintillator that is made lightproof, a scintillator that has almost thesame arrangement as that of a general scintillator but is modified toemit no light, a scintillator made of a heavy metal, or the like isused. Note that the dummy scintillator pieces 62 are positioned/mountedsuch that the x-ray incident surfaces of the scintillator pieces 61 anddummy scintillators 62 become almost flush with each other.

The length of each dummy scintillator 62 in the channel direction may beequal to that of the scintillator block 16. In this case, the length ofa side surface of the dummy scintillator 62 remains unchanged, but thelength of the x-ray incident surface of the dummy scintillator 62 in theslice direction is equal to the length of the 61 in the slice direction,and the length in the channel direction is equal to the length of thescintillator block 16.

In this case, since the respective scintillator pieces and photodiodeshave almost the same shape and size, the respective scintillator piecesand photodiodes are likely to exhibit the same x-ray conversion ratio.However, since no photodiodes are used for the dummy scintillators 62,no data can be acquired from the dummy scintillators 62. If, therefore,a plurality of element blocks 15 are arrayed in the slice direction A,data acquisition omission portions are present between the elementblocks. In this case, therefore, the data precision can be improved byperforming weighted interporation such that omitted data is obtained byaveraging data acquired by photodiodes adjacent to each data acquisitionomission portion in the slice direction A or photodiodes adjacent to theadjacent photodiodes in the channel direction C. The range of data andweights used for interpolation are set in consideration of the purposeof inspection, the precision of data obtained by the elements at theends, expected resolution, and the like.

As described above, the method using dummy scintillators can bepracticed by only adding shielding means to a conventional scintillatorblock, and hence is very versatile.

FIG. 18 is a circuit diagram of a portion of the element block 15. Theelement block 15 has a plurality of photodiodes 17 arranged in the formof an m×n matrix. Signal lines 74 are connected to the outputs of thephotodiodes 17 through a plurality of transistor switches 72. Theoutputs of m photodiodes 17 arrayed in a line along the slice directionA are commonly connected to the same signal line 74. The gates of ntransistor switches 72 arrayed in the channel direction C are commonlyconnected to the same control line 75.

When an x-ray beam strikes a given scintillator piece 11, the x-ray beamis converted into light by the scintillator piece 11. This light isconverted into an electrical signal by the corresponding photodiode 17.While the transistor switch 72 is off, charges are stored in thephotodiode 17. A plurality of control lines 75 are sequentiallyactivated. A plurality of switches 72 are sequentially tuned on insynchronism with the above operation. A plurality of switches 75 aresequentially tuned on in the slice deirection A and tuned on in thechannel deirection C in a parallel. As a consequence, pieces of chargeinformation in a plurality of slices are serially read out. In the priorart, one signal line is connected to each photodiode. If, however, aplurality of photodiodes in a slice line in each channel are commonlyconnected to a signal line, the number of signal lines can be greatlyreduced.

When one slice is to be constituted by a predetermined number ofadjacent photodiodes, analog signal addition can be implemented bysimultaneously turning on the switches 72 of connected to the adjacentcontrol line 75. Thereby data partially added in a slice can be output.

FIG. 19 shows another arrangement of a portion of the element block 15.An integrator 76 is interposed between the photodiode 17 and thetransistor switch 72. As each integrator 76, a type of integrator havingan amplifier connected in parallel with a capacitor or another type ofintegrator is used.

Since an output from each photodiode is an analog current signal. Toperform signal processing for such a signal in a general computer, thiscurrent signal is converted into a voltage signal, and the voltagesignal is converted into a digital signal. In the case shown in FIG. 19,the integrator 76 between the photodiode 17 and the transistor switch 72performs current/voltage conversion. This eliminates the necessity toprovide any current/voltage conversion circuit for the data acquisitioncircuit board 25. In addition, the response speed increases.Furthermore, since the path of an output from the amplifier of theintegrator 76 elongates, and the path of data input to the amplifierwhich is susceptible to disturbances such as noise shortens, resultingin an increase in resistance to disturbances such as noise.

As shown in FIG. 20, a control signal generating circuit 77 is formed ona corner of the element block 15. Referring to FIG. 20, an illustrationof an scintillator in the direction of the drawing surface is omitted,only a portion of the scintillator is indicated by the hatching toexplain the positional relationship between the scintillator and otherelements. The switch 72 and control signal generating circuit 77 arehidden behind the scintillator when viewed from the x-ray tube 131,thereby preventing a malfunction and damage due to radiation of x-rays.Forming the control signal generating circuit 77 on the corner of theelement block 15 eliminates the necessity to form a plurality ofinterconnections for supplying control signals from the outside of theelement block to a plurality of control signals 75. Since only a fewcontrol signal is required to be supplied from the outside of theelement block to the control signal generating circuit 77, thearrangement of interconnections can be simplified.

Signal read operation according to an embodiment of the presentinvention will be described next. FIG. 21 is a schematic circuit diagramshowing one detector module of the radiation detector 127 and a portionof the data acquisition circuit 134 which corresponds to one module. Asdescribed above, one detector module 34 has four element blocks 15-1,15-2, 15-3, and 15-4 arrayed in the slice direction. Assume that in eachof the element blocks 15-1, 15-2, 15-3, and 15-4, a plurality ofdetection elements 42, each constituted by the scintillator piece 11 andphotodiode 17, are arranged in the form of a 24×64 matrix.

In each of the element blocks 15-1, 15-2, 15-3, and 15-4, 24 signallines 74 and 64 control lines are arranged in columns and rows, and thedetection elements 42 are respectively arranged on the intersections ofthe lines. The outputs of the photodiodes 17 of the 64 detectionelements 42 arrayed in a slice line in each channel are connected to thecommon signal lines 74 through the 64 transistor switches. These signallines 74 are connected to each other between the element blocks. The 24signal lines 74 are connected to each amplifier 44. The gates of 24element transistors arrayed in a channel line in each slice are commonlyconnected to the 64 control lines 75.

A vertical shift-register 40-1, 40-2, 40-3 and 40-4 sequentiallysupplies pulses to the 64×4 control lines 75 across the four elementblocks 15-1, 15-2, 15-3, and 15-4. With this operation, as shown in FIG.22, signals are sequentially read out from the 64×4 detection elements42 arranged in a slice line in each channel to the amplifier 44converted into voltage signals by an amplifier 44 connected to theoutput line 47, and are further converted into digital signals by ananalog/digital converter (ADC) 46. This operation is executed the firstsignal line 74—the 24th signal line 74 in parallel. Such signal readoperation in the 38 detector modules 34 are executed in parallel.

FIG. 23 shows another arrangement of the detector module 34. In thiscase, the signal lines 74 are not connected between the element blocks,and output bus lines 47-1, 47-2, 47-3, and 47-4, and amplifiers 44-1,44-2, 44-3, and 44-4 are respectively provided for the element blocks15-1, 15-2, 15-3, and 15-4. Outputs from the amplifiers 44-1, 44-2,44-3, and 44-4 are output through switches 41-1, 41-2, 41-3 and 41-4 andthe common analog/digital converter 46. The switches 41-1, 41-2, 41-3and 41-4 are sequentially operated.

A readout pulses for amplifiers 44-1, 44-2, 44-3 and 44-4 sequentiallysupplies to these being shifted from each other by ¼ the time of a dataperiod (1/fc). With this operation, as shown in FIG. 24, signal readoperation is performed in accordance with the interleaving scheme. Morespecifically, signal reads of the photodiodes 17 of the three elementblocks 15-2, 15-3, and 15-4 are interleaved between a signal read of agiven photodiode 17 of the element block 15-1 and a signal read of theadjacent photodiode 17 in the slice direction A. This scheme can realizehigh-speed read operation.

In addition, signals may be parallelly read out from the four elementblocks 15-1, 15-2, 15-3, and 15-4 in one detector module 34 byrespectively providing analog/digital converters 46-1, 46-2, 46-3, and46-4 for the element blocks 15-1, 15-2, 15-3, and 15-4.

(Second Embodiment)

This embodiment relates to an x-ray CT apparatus (x-ray computedtomography apparatus; CT scanner) equipped with the 2D array typeradiation detector having a large field of view according to the firstembodiment. Note that x-ray CT apparatuses include various types, e.g.,a rotate/rotate type which an x-ray tube and radiation detectorintegrally rotate around an object, and a stationary/rotate type inwhich many detection elements are arrayed in the form of a ring, andonly the x-ray tube rotates around an object. The present invention canbe applied to any type. This embodiment will be described below as arotate/rotate type of apparatus that has currently become mainstream. Toreconstruct 1-volume voxel data (or one tomographic image), projectiondata corresponding to one rotation about the object, i.e., about 360°,is required. In a half-scan method, projection data corresponding toabout 210 to 240° is required. The present invention can be applied toeither of these schemes. Assume that 1-volume voxel data (or one sheetof a tomographic image) is reconstructed from projection datacorresponding to about 360° as in the former general scheme.

FIG. 26 shows the arrangement of an x-ray CT apparatus according to thisembodiment. FIG. 27 is a perspective view of the radiation detector inFIG. 26. A rotating ring 102 is rotated at a speed as high as onerotation per sec by a gantry driving unit 107. An x-ray tube 101 forgenerating an x-ray cone beam (rectangular pyramid) to an object Pplaced in an effective field-of-view region FOV is mounted on therotating ring 102. A high-voltage generator 109 supplies power requiredfor the radiation of x-rays to the x-ray tube 101 through a slip ring108.

A radiation detector 103 for detecting x-rays transmitted through theobject P is attached to the rotating ring 102 in a direction to opposethe x-ray tube 101. In the radiation detector 103, a plurality ofdetection elements, each constituted by a pair of scintillator piece andphotodiode as described in the first embodiment, are arranged in theform of a matrix in the slice direction of the object and the channeldirection perpendicular to the slice direction. For example, severalthousand detection elements are arranged in the channel direction,whereas several hundred detection elements are densely arranged in theslice direction.

Enormous data about all the M×N channels detected by the radiationdetector 103 (M×N channel data per view will be referred to as “2Dprojection data” hereinafter) are temporarily collected by a dataacquisition circuit (DAS) 104 and transmitted altogether to the dataprocessing unit on the stationary side through a noncontact-type datatransmitting unit 105 using optical communication. Detecting operationby the radiation detector 103 is repeated e.g., about 1,000 times duringone rotation (about one sec) to generate enormous 2D projection datacorresponding to M×N channels 1,000 times per sec (rotation). Totransmit such enormous 2D projection data, which are generated at highspeed, without any time delay, the data acquisition circuit 104 andnoncontact-type data transmitting unit 105 are designed to performultra-high speed processing.

The following components are mutually connected to the data processingunit through a data/control bus 300: a host controller 110 serving as amain unit, a pre-processing unit 106 for performing pre-processing suchas data correction, a storing unit 111, a secondary storing unit 112, adata processing unit 113, a reconstructing unit 114, an input device115, and a display unit 116. In addition, an external image processingunit 200 made up of a secondary storage unit 201, data processing unit202, reconstructing unit 203, input device 204, and display unit 205 isconnected to the data processing unit through the data/control bus 300.

FIG. 28 shows data processing and its flow. An x-ray beam transmittedthrough the object is converted into 2D projection data of an analogelectrical signal by the radiation detector 103 and further convertedinto 2D projection data of a digital electrical signal by the dataacquisition circuit 104. Thereafter, the data is sent through thenoncontact-type data transmitting unit 105 to the pre-processing unit106 for performing various correction operations. The 360° 2D projectiondata, i.e., 1,000 sets of 2D projection data, which have undergonesensitivity correction, x-ray intensity correction, and the like in thepre-processing unit 106 are sent to the reconstructing unit 114 directlyor after temporarily stored in the storing unit 111. These data are thenreconstructed into x-ray absorption coefficient 3D distribution data (tobe referred to as “volume data (collection of voxel data)”) in a widetarget region (volume) in the slice direction according to a 3D imagereconstruction algorithm represented by, for example, a so-calledFeldkamp method. This 3D distribution data is typically reconstructed asa collection of multislice tomographic image data.

The reconstructed volume data is sent to the data processing unit 113directly or after temporarily stored in the storing unit 111. This datais then converted into so-called pseudo-3D image data, e.g., atomographic image of an arbitrary slice, a projection image from anarbitrary direction, or a 3D surface image of a specific organ which isobtained by rendering processing, in accordance with an instruction froman operator, and is displayed on the display unit 116.

Although data processing such as reconstruction and slice conversion anddisplay operation are generally performed within an x-ray CT apparatus100, these operations may be executed by the external image processingunit 200. When the external image processing unit 200 is to be used,data sent from the x-ray CT apparatus 100 to the image processing unit200 does not interfere with the effects of this embodiment regardless ofthe state of the data, i.e., a state before reconstruction, a stateafter reconstruction, or a state immediately before display operationafter data processing.

Although the voxel size of the above volume data changes depending onthe size of one detection element of the radiation detector 103, thegeometry of the system, the data acquisition speed, and the like, theminimum voxel size should be about 0.5 mm×0.5 mm×0.5 mm. The apparatus100 equipping the detector of the first embodiment can acquire big sizeand isotropic volume data in one rotation. Further the voxel data issuccessively acquired in a wide region. Therefore, a resolution can befixed between tomographic images for cross sections. This is advantageto a clinical diagnosis.

The operator of the system selects and sets one of the display formsdescried above, i.e., a tomographic image of an arbitrary slice, aprojection image from an arbitrary direction, and 3D surface display,which have already been widely practiced, in accordance with thepurposes of an inspection and diagnosis. Images in different forms aregenerated from one volume data and displayed. Display modes include amode of simultaneously displaying a plurality of types of images as wellas a mode of displaying only one type of image. The operator can switchthese modes in accordance with a purpose.

As shown in FIG. 29, in addition to a tomographic image of a slice(axial slice) perpendicular to the body axis which is obtained byconventional x-ray CT apparatus, tomographic images of arbitrary slicesinclude tomographic images of slices perpendicular to the axial slice,e.g., a saggital plane and coronal plane, and tomographic images ofslices oblique to these slices. Voxel data of a designated slice with adesignated thickness are extracted from the above volume data anddisplayed altogether. A projection image from an arbitrary direction isused to display, for example, the maximum value and cumulative value ofvoxel data arranged in a set direction as a 2D image with respect to thevolume data. 3D surface display is a method of extracting a surface witha set threshold and displaying the surface as a 3D image by shadingbased on a set light source. With this method, the operator can grasp aninternal structure by observing while changing the threshold.

In 1-rotation scanning, by performing the above data processing, onevolume data about a region of interest as wide as 30 cm in the slicedirection can be obtained, without any time difference in the slicedirection, from 2D projection data from many directions which areobtained by only one rotation. The operator can observe a tomographicimage at a given time other than a tomographic image of an axial slice.

When the same processing as that in 1-rotation scanning is to berepeatedly performed for 2D projection data from many directionsobtained by a plurality of rotations in continuous rotation scanning, aplurality of volume data are obtained instead of one volume data. Evenif reconstruction is performed every rotation, data sets equal to thenumber of rotations can be obtained. In addition, by shifting the range(rotational angle range of the system) of data used for reconstructionlittle by little, many volume data that slightly differ in time can beobtained.

As in the case of 1-rotation scanning, as a display image form, one ofthe following forms: a tomographic image of an arbitrary slice, aprojection image from an arbitrary direction, and 3D surface display,can be selected in accordance with the settings in the system which aremade by the operator.

Images that slightly differ in time are generated in a set form from theabove volume data that slightly differ in time, and sequentiallydisplayed, as shown in FIG. 30. This allows the operator to observe theimages in the set form in real time as moving images. Operation ofdisplaying images as moving images concurrently with this continuousscanning will be referred to as CT fluoroscopy.

FIG. 31 shows the temporal flow from scanning in this CT fluoroscopy toimage displaying on one time scale. Assume that the angular range ofprojection data required to reconstruct one 3D image data is 360°.Obviously, this range may be set to 180°+view angle. First of all, thex-ray tube 101 and radiation detector 103 continuously rotate around theobject at high speed. The time required for one rotation is representedby to. Projection data that are sequentially acquired are subjected topre-processing almost in real time. The reconstructing unit 114 thenreconstruct 3D image data “I” on the basis of the 360° projection datahaving undergone the pre-processing. The data processing unit 113generates image data “DI” of a tomographic image of an arbitrary slice,a projection image from an arbitrary direction, 3D surface image, or thelike on the basis of the reconstructed 3D image data “I”. This imagedata “DI” is displayed on the display unit 116.

In CT fluoroscopy, a series of operations from scanning to imagedisplaying are concurrently performed, and images are sequentiallyreconstructed while continuous scanning is performed. These images aresequentially displayed to be displayed as moving images.

To implement this CT fluoroscopy, the reconstructing unit 114 has theperformance required to reconstruct the 3D image data I within a timeshorter than the time t0 required to acquire projection datacorresponding to a predetermined angular range (360° in this case)concurrently with acquisition operation of projection data (scanning).The data processing unit 113 has the performance required to generatethe display image data DI from 3D image data within a time shorter thanthe reconstruction time for the 3D image data I. The display unit 116has a counter, memory, and the like which are required to startdisplaying the image data DI a predetermined-time after a start point Tsor end point Te of an interval of acquisition operation of projectiondata from which the image data DI originates.

To facilitate observation of images as moving images, this apparatusfurther includes the following means.

(1) Displaying can be performed not only in the forward direction butalso in the reverse direction (reverse playback).

(2) An automatic-updating mode or manual updating mode can be selectedas an image updating (switching) mode, and image switching can be doneeven during display operation.

(3) In the automatic updating mode, the operator designates a startpoint (moving image playback start point) and end point (moving imageplayback end point), and image updating is done at a predeterminedupdating speed (image switching speed (moving image playback speed)).

(i) The start and end points can be changed even during displayoperation.

(ii) The predetermined updating speeds include the following modes:

(a) actual time intervals based on the scanning speed and reconstructionintervals;

(b) slow display

(c) frame display

(d) fast (double-speed) display

(iii) In addition to preset speeds, displaying is performed at anarbitrary speed set by the operator.

(iv) Updating speed can be changed even during display operation.

(v) When displaying is done up to the end point, displaying is repeatedfrom the start point.

(4) In the manual updating mode, updating is performed in accordancewith the operation performed by the operator.

To easily grasp the relationship between overall movement and an imagethat is being displayed, all or some of images in the overall time rangecan be displayed as index images concurrently with a main image.

(1) The time of a main image is displayed on an index image. The time ofthe main image is set on the index image and the playback start point ofa moving image is switched to another point.

(2) Index images can be obtained by reducing images or decreasing theresolution of images in the data processing unit 113, and a plurality ofindex images are simultaneously displayed in one window as a list.

(3) Index images are not generated and displayed with respect to allimages as targets, but a plurality of images in the playback period areproperly thinned out and selected.

(i) Images are thinned out and displayed at predetermined timeintervals.

(ii) A portion corresponding to a fast motion between images isextracted and displayed.

(4) Index images are used to display the time zones before and after amain image and updated as the main image is updated.

Information that changes with time, e.g., the CT value of an ROI orelectrocardiogram, is displayed in the form of a graph, concurrentlywith main image displaying, and the time of the main image is alsodisplayed on the graph that is being displayed.

Additional advantages and modifications will readily occur to thoseskilled in the art. Therefore, the invention in its broader aspects isnot limited to the specific details and representative embodiments shownand described herein. Accordingly, various modifications may be madewithout departing from the spirit or scope of the general inventiveconcept as defined by the appended claims and their equivalents.

What is claimed is:
 1. An x-ray CT apparatus, comprising: a gantryincluding: an x-ray source which generates an x-ray cone beam, aradiation detector having a plurality of element blocks, each elementblock having a plurality of detection elements which are arranged in amatrix form and which detect x-rays transmitted-through an object, and arotating device which allows the x-ray source and the radiation detectorto rotate continuously to permit said gantry to continuously acquireprojection data on the object; and a reconstructing unit whichreconstructs volume data acquired at an arbitrary timing, on the basisof the projection data, wherein each of said plurality of element blocksis formed on a corresponding one of a plurality of block substrates,said plurality of block substrates are formed in a direction parallel toa rotational axis of the radiation detector to obtain each of aplurality of modules, and said plurality of modules are arranged in adirection perpendicular to the rotational axis of the radiation detectoron a base of the radiation detector.
 2. An x-ray CT apparatus accordingto claim 1, wherein said gantry acquires projection data itemscorresponding to a plurality of timings; and said reconstructing unitreconstructs volume data items corresponding to the plurality oftimings.
 3. An x-ray CT apparatus according to claim 1, wherein saidvolume data is a collection of isotropic voxel data.
 4. An x-ray CTapparatus according to claim 1, wherein said plurality of element blocksare each arranged on block substrates further arranged on a plurality ofmodule bases that are included in a plurality of detector modules thatare arrange on a base of the radiation detector.
 5. An x-ray CTapparatus according to claim 1, wherein the element blocks of theradiation detector are arranged in such a manner as to form a polygon asviewed in a direction perpendicular to the rotational axis of theradiation detector, and are arranged substantially in a plane as viewedin a direction parallel to the rotational axis of the radiationdetector.
 6. An x-ray CT apparatus according to claim 1, wherein theelement blocks of the radiation detector are arranged in such a manneras to form an arc, with a focal point of the x-ray source as a center,as viewed in a direction perpendicular to the rotational axis of theradiation detector.
 7. An x-ray CT apparatus according to claim 5,wherein each of said element blocks includes: a semiconductor substrate;a plurality of photodiodes formed on the semiconductor substrate andarranged in a matrix pattern; and a scintillator block provided on thephotodiodes.
 8. An x-ray CT apparatus according to claim 4, wherein eachof the detector modules includes a data collecting circuit whichcollects signals from the detection elements.
 9. An x-ray CT apparatusaccording to claim 8, wherein said data collecting circuit includes anamplifier and an analog/digital converter.
 10. An x-ray CT apparatusaccording to claim 4, wherein each of said plurality of detector modulesis detachably provided on the detector base.
 11. An x-ray CT apparatusaccording to claim 7, wherein said scintillator block includes ascintillator niece located, at an end as viewed in the directionparallel to the rotational axis of the radiation detector, and thescintillator piece includes a space which arranges wiring connecting thephotodiodes to the substrate.
 12. An x-ray CT apparatus according toclaim 4, wherein said detector modules are attached to correspondingcollimator modules.
 13. An x-ray CT apparatus according to claim 12,wherein said collimator modules include collimator plates, and thedetection elements of the detector modules are greater in number thanthe collimator plates.
 14. An x-ray CT apparatus according to claim 13,wherein said collimator modules include a first collimator module and asecond collimator module which share a predetermined collimator plate.15. An x-ray CT apparatus according to claim 7, wherein said photodiodesare arranged in a matrix form of m×n, and comprise m photodiodes whichare arrayed in a body axis direction of the object and which includeoutput terminals connected to a common signal line, and n switchelements gates arrayed in the direction perpendicular to the rotationalaxis of the radiation detector and connected to a common control line.16. An x-ray CT apparatus according to claim 15, wherein said signalline is connected to the output terminals at a position between theelement blocks of the detector modules.
 17. An x-ray CT apparatusaccording to claim 15, further comprising a read control circuitprovided on the substrate, said read control circuit controlling anON/OFF operation of the switching elements and reading out signals fromthe detection elements.
 18. An array CT apparatus according to claim 17,wherein said read control circuit sequentially turns on the switchingelements along the signal line with respect to the plurality of elementblocks in said detector module.
 19. An x-ray CT apparatus according toclaim 17, wherein said read control circuit sequentially turns on theswitching elements along the signal line, and executes the sequentialoperation between the plurality of element blocks in said detectormodule with a predetermined time shift such that signal reads of elementblocks in said detector module are interleaved in an interval of asignal read of the a given element block in said detector module.
 20. Anx-ray CT apparatus according to claim 17, wherein said read controlcircuit executes operations of sequentially turning on the switchingelements along the signal line in parallel with respect to the pluralityof element blocks in said detector module.
 21. An x-ray CT apparatusaccording to claim 1, wherein the plurality of detection elements arearranged parallel and perpendicular to the rotational axis of theradiation detector.